Method and Apparatus to Produce Ultrasonic Images Using Multiple Apertures

ABSTRACT

A combination of an ultrasonic scanner and an omnidirectional receive transducer for producing a two-dimensional image from received echoes is described. Two-dimensional images with different noise components can be constructed from the echoes received by additional transducers. These can be combined to produce images with better signal to noise ratios and lateral resolution. Also disclosed is a method based on information content to compensate for the different delays for different paths through intervening tissue is described. The disclosed techniques have broad application in medical imaging but are ideally suited to multi-aperture cardiac imaging using two or more intercostal spaces. Since lateral resolution is determined primarily by the aperture defined by the end elements, it is not necessary to fill the entire aperture with equally spaced elements. Multiple slices using these methods can be combined to form three-dimensional images.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.13/632,929, filed Oct. 1, 2012; which application is a continuation ofU.S. patent application Ser. No. 13/215,966, filed Aug. 23, 2011, nowU.S. Pat. No. 8,277,383; which application is a continuation of U.S.patent application Ser. No. 11/865,501, filed Oct. 1, 2007, now U.S.Pat. No. 8,007,439; which application claims the benefit of U.S.Provisional Patent Applications No. 60/862,951, filed Oct. 25, 2006, andNo. 60/940,261, filed May 25, 2007; all of which are incorporated byreference in their entirety herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates generally to imaging techniques used inmedicine, and more particularly to medical ultrasound, and still moreparticularly to an apparatus for producing ultrasonic images usingmultiple apertures.

2. Discussion of Related Art Including Information Disclosed Under 37CFR §§1.97, 1.98

In conventional ultrasonic imaging, a focused beam of ultrasound energyis transmitted into body tissues to be examined and the returned echoesare detected and plotted to form an image. In echocardiography the beamis usually stepped in increments of angle from a center probe position,and the echoes are plotted along lines representing the paths of thetransmitted beams. In abdominal ultrasonography the beam is usuallystepped laterally, generating parallel beam paths, and the returnedechoes are plotted along parallel lines representing these paths. Thefollowing description will relate to the angular scanning technique forechocardiography (commonly referred to as a sector scan). However, thesame concept with modifications can be implemented in abdominalscanners.

The basic principles of conventional ultrasonic imaging are welldescribed in the first chapter of Echocardiography, by Harvey Feigenbaum(Lippincott Williams & Wilkins, 5^(th) ed., Philadelphia, 1993). Thesewill not be repeated here except as necessary to illustrate thedifferences between the conventional techniques and the presentinvention.

It is well known that the average velocity v of ultrasound in humantissue is about 1540 m/sec, the range in soft tissue being 1440 to 1670m/sec (see for example P. N. T. Wells, Biomedical Ultrasonics, AcademicPress, London, New York, San Francisco, 1977). Therefore, the depth ofan impedance discontinuity generating an echo can be estimated as theround-trip time for the echo multiplied by v/2, and the amplitude isplotted at that depth along a line representing the path of the beam.After this has been done for all echoes along all beam paths, an imageis formed, such as the image 10 shown in FIG. 1, in which a circle hasbeen imaged. The gaps between the scan lines are typically filled in byinterpolation. One of the earliest interpolation algorithms applied toechocardiography was described in U.S. Pat. No. 4,271,842, to Specht etal.

In order to insonify the body tissues, a beam formed either by a phasedarray or a shaped transducer is scanned over the tissues to be examined.Traditionally, the same transducer or array is used to detect thereturning echoes. This design configuration lies at the heart of one ofthe most significant limitations in the use of ultrasonic imaging formedical purposes; namely, poor lateral resolution. Theoretically thelateral resolution could be improved by increasing the aperture of theultrasonic probe, but the practical problems involved with aperture sizeincrease have kept apertures small and lateral resolution large.Unquestionably, ultrasonic imaging has been very useful even with thislimitation, but it could be more effective with better resolution.

In the practice of cardiology, for example, the limitation on singleaperture size is dictated by the space between the ribs (the intercostalspaces). For scanners intended for abdominal and other use, thelimitation on aperture size is not so obvious, but it is a seriouslimitation nevertheless. The problem is that it is difficult to keep theelements of a large aperture array in phase because the speed ofultrasound transmission varies with the type of tissue between the probeand the area of interest. According to the book by Wells (cited above),the speed varies up to plus or minus 10% within the soft tissues. Whenthe aperture is kept small, the intervening tissue is, to a first orderof approximation, all the same and any variation is ignored. When thesize of the aperture is increased to improve the lateral resolution, theadditional elements of a phased array may be out of phase and mayactually degrade the image rather than improving it. The instantdisclosure teaches methods to maintain all of the information from anextended phased array “in phase” and thus to achieve sought-afterimproved lateral resolution.

In the case of cardiology, it has long been thought that extending thephased array into a second or third intercostal space would improve thelateral resolution, but this idea has met with two problems. First,elements over the ribs have to be eliminated, leaving a sparsely filledarray. New theory is necessary to steer the beam emanating from such anarray. Second, the tissue speed variation described above, but notadequately addressed until this time, needs to be compensated. The samesolution taught in this disclosure is equally applicable formulti-aperture cardiac scanning, or for extended sparsely populatedapertures for scans on other parts of the body.

BRIEF SUMMARY OF THE INVENTION

The present invention solves both the problem of using more than oneintercostal space and the problem of accommodating unknown phase delaysfrom using elements spread over a large sparse aperture. The solutioninvolves separating the insonifying probe from the imaging elements. Theseparation can be a physical separation or simply a separation inconcept wherein some of the elements of the array can be shared for thetwo functions.

A single omni-directional receive element (such as a receive transducer)can gather all of the information necessary to reproduce atwo-dimensional section of the body. Each time a pulse of ultrasoundenergy is transmitted along a particular path, the signal received bythe omnidirectional probe can be recorded into a line of memory. [Theterms “omni-directional probe,” “omni probe” and/or “omni,” are usedsynonymously herein to mean an omnidirectional probe.] When this is donefor all of the lines in a sector scan, the memory can be used toreconstruct the image. This can be accomplished in the same time as datais being collected for the next frame.

There are numerous advantages to this approach, and these comprise theobjects and advantages of the present invention. They include, amongothers:

The dominance of specular reflections so prominent in reconstructingimages by returns to the main probe is greatly attenuated.

More than one omnidirectional probe can be used. Each one can be used toreconstruct an entire sector image but with different point spreadfunctions. These can be combined to produce an image with a sharperpoint spread function.

Compensations can be made for different delays in different pathsthrough the tissue.

Many more scan lines can be reconstructed than the number of pulsesgenerated by the main probe. This overcomes the traditional limit of thenumber of scan lines by the speed of ultrasound in tissue, tissue depthof interest, and the time allowed between frames, which is typically1/30^(th) second.

Artificial scan lines can be considered as overlapping, and each pixelon an output image can be imaged from information from more than oneomni line of data. Therefore the output pixel can be averaged frommultiple data, thus improving the signal-to-noise ratio.

Omnidirectional probes can be placed in multiple intercostal spaces, thesuprasternal notch, the substernal window, multiple apertures along theabdomen and other parts of the body, and even on the end of a catheter.An advantage in using omnidirectional probes for catheter placement isthat no steering is required of the probe.

Probes can be placed either on the image plane, off of it, or anycombination. When placed away from the image plane, omni probeinformation can be used to narrow the thickness of the sector scanned.

There has thus been broadly outlined the more important features of theinvention in order that the detailed description that follows may bebetter understood, and in order that the present contribution to the artmay be better appreciated. Additional objects, advantages and novelfeatures of the invention will be set forth in part in the descriptionas follows, and in part will become apparent to those skilled in the artupon examination of the following. Furthermore, such objects, advantagesand features may be learned by practice of the invention, or may berealized and attained by means of the instrumentalities and combinationsparticularly pointed out in the appended claims.

Still other objects and advantages of the present invention will becomereadily apparent to those skilled in this art from the followingdetailed description, which shows and describes only the preferredembodiments of the invention, simply by way of illustration of the bestmode now contemplated of carrying out the invention. As will berealized, the invention is capable of modification in various obviousrespects without departing from the invention. Accordingly, the drawingsand description of the preferred embodiment are to be regarded asillustrative in nature, and not as restrictive.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS

FIG. 1 is a diagrammatic view of a simulation showing a circular objectimaged by a conventional sector scanner.

FIG. 2 is a schematic diagram of the axes representing the relativepositions of the insonifying and omni-directional probes.

FIG. 3 is a graph showing the orientation of the point spread functionof an imaging system using a single omnidirectional probe as a functionof the position of the omnidirectional probe.

FIG. 4 is a graph showing the orientation of the point spread functionof an imaging system using a single omnidirectional probe as a functionof the position of the omnidirectional probe when it is not in the scanplane.

FIG. 5 is a diagrammatic view of a simulation showing the same circularobject of FIG. 1 imaged by data received by a single omni elementlocated at x₀=40 mm, y₀=0 mm, z₀=0 mm.

FIG. 6 is a schematic diagram illustrating the relative positions ofprobes showing, in addition, points A-A′, which have equal round tripdistances and times.

FIG. 7 is a schematic diagram showing a possible fixture for positioningan omni-directional probe relative to the main (insonifying) probe.

FIG. 8 is a schematic diagram showing a non-instrumented linkage for twoprobes.

FIG. 9 is a schematic diagram showing variables for computation of x andy positions from received echoes.

FIG. 10A is a phantom image taken with a standard Acuson 128 XP-10 witha 3.5 MHz transducer and harmonic processing.

FIG. 10B is the same phantom image as that shown in FIG. 10A and takenwith the same XP-10, wherein the center 64 elements were obscured butexternal processing employed to show improved lateral resolution. Theprogressions of anechoic areas on the phantom are 8 mm diameter, 6 mm, 4mm, 3 mm, and 2 mm.

FIG. 11 is an image of the same phantom produced by the same transduceras the images in FIGS. 10A and 10B, with the center obscured, but withsubstantial image processing over multiple scans. Note that even thoughthe total aperture is only 19 mm that the 2 mm diameter anechoic areasare now visible. Lateral resolution could be greatly improved if the twoparts of the transducer were physically separated and the phased delaysreprogrammed for the resulting geometry.

FIG. 12A is a schematic perspective view showing an adjustable,extendable hand held two-aperture probe (especially adapted for use incardiology US imaging). This view shows the probe in a partiallyextended configuration.

FIG. 12B is a side view in elevation thereof showing the probe in acollapsed configuration.

FIG. 12C shows the probe extended so as to place the heads at a maximumseparation distance permitted under the probe design, and poised forpushing the separated probe apertures into a collapsed configuration.

FIG. 12D is a side view in elevation again showing the probe in acollapsed configuration, with adjustment means shown (i.e., as scrollwheel).

FIG. 12E is a detailed perspective view showing surface features at thegripping portion of the probe.

FIG. 13 is a 3D image highlighting the anechoic tubes of the ATS Model539 phantom.

DETAILED DESCRIPTION OF THE INVENTION

A key element of the present invention is that returned echoes inultrasonography can be detected by a separate relatively non-directionalreceive transducer located away from the insonifying probe (transmittransducer), and the non-directional receive transducer can be placed ina different acoustic window from the insonifying probe. This probe willbe called an omni-directional probe because it can be designed to besensitive to a wide field of view.

If the echoes detected at the omni probe are stored separately for everypulse from the insonifying transducer, it is surprising to note that theentire two-dimensional image can be formed from the information receivedby the one omni. Additional copies of the image can be formed byadditional omni-directional probes collecting data from the same set ofinsonifying pulses.

A large amount of straightforward computation is required to plot theamplitude of echoes received from the omni. Referring now to FIG. 2, inwhich there is shown the position of the omni-directional probe 100relative to the position of the insonifying (main) probe 120 and theinsonifying beam 130. The position of the omni-directional proberelative to the beam is indicated by x₀, y₀ and z₀ 140, where x₀ and y₀are in the scan plane 150 scanned by the insonifying beam and z₀ isdistance perpendicular to that plane. Instead of simply plotting thedepth along the scan line d=tv/2 (where t is the round-trip time, it isnow computed as that point d=sqrt(x̂2+ŷ2) for whichtv=d+sqrt((x−x₀)̂2+(y−y₀)̂2+z₀̂2).

This procedure will produce a sector scan image similar to that usingthe conventional technique except that the point spread function will berotated. FIG. 3 shows the orientation 200 of the psf as a function ofthe position of the omni probe. A single point at x=0, y=70 mm, z=0 isthe point being imaged, and the groups of dots 210 each indicate theorientation of the psf if the omni-probe were placed at the location ofthe center of each group. The insonifying probe (main probe) is locatedat the center group 220 on the bottom of the figure. In this simulationthe horizontal (x axis) 230 shown goes from −40 mm to +40 mm. Thevertical (y axis) 240 goes from 0 to 80 mm depth.

FIG. 4 shows the orientation 300 of the psf as a function of theposition of the omni probe if it is not in the scan plane. In thissimulation the horizontal (x axis) 310 shown again goes from −40 mm to+40 mm, but the vertical axis 320 is the z axis (distance away from thescan plane) and goes from 0 to 80 mm away from the scan plane.

FIG. 5 shows a plot of the same circular object 20 as in FIG. 1,plotted, however, from data received by a single omni element. Othercomplete two dimensional reconstructions may be formed using data fromadditional omni elements, if desired.

Specular reflection is reduced using the omni probe compared to usingthe main probe for both insonification and detection. This is becauseall parts of a surface normal to the main beam are insonified with thesame phase. When the phase is such that maximum echo is returned, allthe echoes add to produce a specular echo. When the signals arenormalized to accommodate the dynamic response of a particular displaydevice, non-specular echoes will tend to drop out.

In contrast to this, and referring now to FIG. 6, there is also shownthe relative positions of probes, but there is also shown points A-A′400, which have equal round trip distances and times. Points A-A′ are ona surface normal to the bisector line a₁=a₂ 410, which also have equalround trip distances and times, are not insonified with the same phase,and do not all reflect equally. This attenuation of the specularreflection is particularly important when visualizing circularstructures which frequently have surfaces normal to the main beam.

An algorithm to plot this data on a rectangular grid is: (a) for eachpoint on the x,y grid, convert x,y to depth and angle; (b) then findclosest angle k scanned by the insonifying beam; (c) if it issufficiently close, then convert x,y to distance to the omni; (d)compute time t=(distance to insonifying beam+distance to omni)/v; and(e) plot amplitude recorded by the omni for the k scan at x,y.

However, more information is available and should be used. It ispossible to use the same technique to plot additional scan lines whichwere not explicitly insonified by the insonifying probe. Because of theinherently wide beam width of medical ultrasonic probes, much tissuebetween intentional scan lines is also insonified and returns echoes.Making use of this information is particularly important when capturingmotion (especially in echocardiography) because the number of pulsesthat can be generated is strictly limited by the speed of ultrasound intissue and the scan repetition rate desired.

The reconstructed image will get better as the angle between the mainbeam and the omni gets larger. However it is not necessary to focus anarrow beam on every element of tissue to be imaged as is true if thedata is not stored and then processed before display. The lateralresolution can be reconstructed using a Wiener filter to be much betterthan the beam width if the noise spectrum is low enough. In onesimulation of 2 circles of diameter 2.2 mm and 4.0 mm, both imaged wellenough that the center was clear even though the beam width was 4.4 mmtapered from 1 to 0 by a cosine function. The Wiener filter is describedin the next section.

There are four main sources of noise in ultrasonic imaging: (1) blur dueto array size not wide enough; (2) shot noise; (3) reverberation frombig interfaces; and (4) speckle.

Multiple probes give independent measures of shot noise, but usingclosely spaced elements in the main probe (if it is a phased array) willnot give independent noise for the other three sources. Adding one ormore omni probes will change the look angle, which will thereby changethe speckle pattern and the reverberation pattern. These can be averagedout to lower the noise power spectrum. The Wiener filter can then beemployed to cancel the blur.

Another way to eliminate speckle is to obtain a good sample of it forestimates of the noise spectrum to then be used in the Wiener filter.

De-blurring and de-noising by these techniques using only an externalomni probe or probes will make it possible to visualize small and movingobjects such as the coronary arteries. In such a case medical personnelcould assess the degree of opening in the lumen or patency of bypassgrafts without resort to invasive catheterization techniques.

When combining more than one image such as one from the main probe andanother from one or more omni probes for the purpose of averaging outthe various sources of noise, it is necessary to compensate for thevariation in ultrasound velocity through different paths. Experimentshave shown that small unaccounted errors in path velocity will displacethe reconstructed image in both horizontal and vertical positions. Crosscorrelation techniques should be used to find the displacement with oneimage taken as reference before addition or other combination of images.

Two possibilities exist with regard to the Wiener filter. In one, aWiener filter can be used separately on each image and then combinethem. Alternatively, one can first combine the images (yielding a morecomplex point spread function) and then employ Wiener filtering.

In order to perform the indicated computations, it is necessary toeither measure or estimate the position (x₀, y₀, z₀) of the omnirelative to the main probe. Exact knowledge of these positions is notrequired because, as we have seen, the displacement of the image fromthe omni probe is also affected by the variation of the velocity ofultrasound in different types of tissue. For this reason it is necessaryto use cross correlation or some other matching criterion to make afinal correction of the position of the omni-generated image beforecombining with the reference image or images.

Determining the Position of the Omni Probe(s):

There are many ways to determine the position of the omni probe.Referring now to FIG. 7, one way is to provide apparatus 500 forpivotally and/or swivelingly mounting the omni probe 510 on a fixture520 attached to the insonifying probe 530. The fixture preferablyincludes articulated joints 540 with sensors (not shown) to measureangles and distances 550 of the links. FIG. 7 illustrates a simplifiedversion of such a fixture, wherein fixed hinges allow movement of onlyx₀ and y₀.

Another method is to have no mechanical connection between the omniprobe and the main probe (except wires for signals and power). Instead,the omni probe can transmit a signal using radio frequencies, light, ora different frequency ultrasound to triangulation receivers mounted onthe main probe or a separate platform.

A third method again has no mechanical connection between the omni probeand main probe. For this method the omni probe (or probes) can beattached to the patient with tape, and the ultrasonographer canmanipulate the main probe to find the best image without regard topositioning the omni probe(s). As indicated in FIG. 4, a two-dimensionalimage can be formed separately from the echoes received from the omniprobe and from the main probe. By adjusting four variables (x₀, y₀, z₀and D, the average difference in time for ultrasound to go throughdifferent tissue types instead of traveling though idealized tissue ofconstant ultrasound velocity), the images can be made to coincide. Thefour variables can be adjusted iteratively to maximize cross-correlationor another measure of similarity. Standard multi-dimensional searchtechniques that could be used include gradient ascent, second order(Newton's method), simulated annealing, and genetic algorithms. A fifthvariable, the angle of the psf which is necessary for deconvolution, isimplied from the first four variables. Misregistration of the images canbe caused by inaccurate estimation of any of the four variables, butgood registration can be achieved by simply adjusting D which will tendto compensate for errors in estimates of the others.

When the application requires the highest resolution compatible withcapturing motion at a high frame rate, the four variables can beestimated over several frames of information. When the ultrasonographerhas selected a good view angle, the frames can be combined at high rateholding x₀, y₀, z₀ and D constant.

When the application requires the highest possible resolution, data canbe captured (perhaps with EKG gating to capture separate images atsystole and diastole) and the multi-dimensional search to optimizematching can be done more accurately although not in real time. Twoadvantages of this approach is that different values of D can be foundfor systole and diastole, and that different psf's can be used fordeconvolution at different depths in the image.

Determining the Position of the Omni Probe(s) Using Correlation of theScan Line Data Rather than Complete 2D Sectors:

A fourth method for determining the position of the omni probe(s)entails replacing the omni probe or probes with a “semi-omni probe” orprobes. The reason for this is to increase the signal to noise ratio byrestricting the sensitive region of the receive transducer to a planerather than a hemisphere. Because of this restriction it is necessary tohave a mechanical linkage to ensure that both the transmit and receivetransducers are focused on the same plane.

Two probes could be placed in any two acoustic windows. In the case ofechocardiography, one would likely be in the normal parasternal windowwhich typically gives the clearest view of the whole heart. A secondwindow available in most patients is the apical view. Another windowusually available is the subcostal. The two probes could even be placedon either side of the sternum or in parasternal windows in twointercostal spaces.

One probe could be the standard phased array cardiac probe. The second(and third, etc.) would be used as receive only. Theoretically it couldbe omnidirectional, but that would necessarily provide lower signals andtherefore low signal to noise ratios (S/N). A better alternative is touse a probe which is ground to be sensitive to a plane of scan butomnidirectional within that plane. A single piece of PZT would workwell, but to minimize the amount of new design required it is alsopossible to use a second probe head similar to the main probe and thenuse individual elements or small groups combined to act as singleelements. The design goal is to use as many elements as possible tomaximize signal to noise ratio while using few enough to minimize anglesensitivity.

In this embodiment 600 (see FIG. 8), the two probes 610, 620, may belinked together with an articulating mechanical linkage, which ensuresthat the plane of scan of each probe includes the other, but thedistance between them is unconstrained. A slave servomechanism is alsopossible, but the mechanical linkage will be described here.

The procedure is to aim the main probe 620 at the target 630 (e.g.,heart) and position the secondary probe 610 at a second window withmaximum received signal strength. One possibility is that the main probebe positioned for a long axis view with the secondary probe over theapex of the heart. Some slight deviation of the long axis view may benecessary in order to maintain the secondary probe in its most sensitivespot.

The secondary probe would now be held on the patient with a mechanicalhousing which allows a fan or rocking motion. The disadvantage of havingtwo probes in fixed positions on the body is that the plane of scan mustinclude these two points. The only degree of freedom is the angle atwhich the scan plane enters the patient's body. For a conventional 2Dexamination this is a severe limitation, but if the goal is to gatherthree-dimensional information, this is not a limitation. The 3Dinformation is obtained by rocking the main probe back and forth througha sufficient angle so that the entire heart is insonified. The secondaryprobe also rocks back and forth by virtue of the mechanical linagebetween the probes. The instantaneous angle of rocking must bemonitored—perhaps by reference to a gyroscope mounted with the mainprobe. The rocking could be actuated by the hand motion of theultrasound technician, or it could be motorized for a more-uniform anglerate. In an alternative preferred embodiment (for echocardiography), themain probe and an array of omni probes are placed in adjacentintercostal spaces using a mechanism as shown in FIG. 12.

Computer software could be provided such that the 2D slices would fill a3D volume of voxels. After adjacent voxels are filled throughinterpolation, 3D information can be displayed as projections or asslices through the volume at arbitrary orientations.

The need for and one important use of the 3D information is covered inU.S. patent application Ser. No. 11/532,013, now U.S. Pat. No.8,105,239, also by the present inventor, and which application isincorporated in its entirety by reference herein.

Yet another variation on this theme is to have the secondary transducermechanically linked to the primary so that each plane of scan containsthe other transducer (as above), but allow rotation of the main probeabout its own axis. In this case the secondary probe would be allowed tomove on the patient's body (properly prepared with ultrasound gel). Itwould have many elements, and an attached computer would scan them allto find those elements which have the strongest return signal.

Estimating Relative Probe Positions from Reflected Signals:

For image reconstruction it is essential to know the position of thesecondary probe (x, y) relative to the main probe. This has to beevaluated separately for each frame of data because of the motion of thepatient, technician, and/or motorized angle actuator. Since the linkagewill prevent any difference in position (z) perpendicular to the scanplane, only x and y need be assessed.

Note that any tilt of the main probe will change the reference axes sothat x and particularly y will change too.

When a pulse is transmitted from the main probe it insonifies a sequenceof tissues in the path of the beam. The returns from the tissues will bereceived by both the main probe and the secondary, digitized and storedin the computer. Echoes from relatively proximate tissues will bedifferent for the two probes, but echoes from mid- to far range will besimilar. It is possible to use cross correlation to find similar smallpatches in the two stored returns. They will be similar except for thetime delay relative to the launching of the pulse from the main probe.The time delay will be related to the offsets x and y. Values for x andy cannot be determined from one set of time delays, but can bedetermined by solving a set of simultaneous equations from two detectedsimilar returns. These could be different patches of the same pulsereturn or from returns from differently directed main pulses.

Referring now to FIG. 9, if the main probe 700 transmits at angle 90°−a710 relative to its centerline and an identifiable packet of returnsoccurs at time t_(1m) at the main probe and at time t.sub.1s at thesecondary (omni) probe 720, then:

-   -   tissue packet at (x₁ y₁) is received at time t_(1m), and        distance m₁=sqrt(x₁ ²+y₁ ²)    -   tissue packet at (x₂, y₂) is received at time t_(2m), and        distance m₂=sqrt(x₂ ²+y₂ ²)    -   t_(1m) corresponds to time of two trips of distance m₁    -   t_(1m)s=2m₁, where s=speed of ultrasound in same units as        m=approx. 1.54×10⁶ mm/sec

t _(1s) s=m ₁+sqrt((x ₁ −x)²+(y ₁ −y)²).

Similarly,

t _(2m) s=2m ₂

t _(2s) s=m ₂+sqrt((x ₂ −x)²+(y ₂ −y)²)

(t _(1s) s−0.5t _(1m))=(x ₁ −x)²+(y ₁ −y)²

(t _(2s) s−0.5t _(2m) s)²=(x ₂ −x)²+(y ₂ −y)²  (1)

Since Xj, y₁₅ x₂, y₂ and the times are known, one can solve the last twosimultaneous equations for x and y. Similarly, if a z offset between thetwo probes is allowed, x, y, and z can be calculated by solving threesimultaneous equations.

Many more measurements from packet pairs are available. One could make ameasurement on several or every scan line (angle) as measured from themain probe. Then we would have many equations in 2 unknowns which can beused to make more-accurate estimations of the 2 unknowns. Since theseare nonlinear equations, a search technique can be utilized. One way toaccomplish this is to compute error squared over a grid of (x, y) pointsusing the equation:

$\begin{matrix}{E_{2} = {\sum\limits_{i = 1}^{N}\; ( {\sqrt{( {x_{i} - x} )^{2} + ( {y_{i} - y} )^{2}} - t_{is} + {0.5\; t_{im}s}} )^{2}}} & (2)\end{matrix}$

The minimum E² will indicate the minimum squared error estimate of (x,y). The search should be conducted over the expected range of x and y tosave time and to avoid spurious ambiguous minima.

When the z component of the relative position is not constrained to bezero, the comparable error squared equation is:

$E_{2} = {\sum\limits_{i = 1}^{N}\; ( {\sqrt{ {( {x_{i} - x} )^{2} + ( {y_{i} - y} )^{2} + z_{i} - z} )^{2}} - t_{is} + {0.5\; t_{im}s}} )^{2}}$

The minimum E² will indicate the minimum squared error estimate of (x,y, z).

If the speed of sound on the return path to the secondary (omni)transducer is different from s due to different types of tissues beingtraversed, the values of x and y (and z if used) will be different fromthe geometric values. However, use of these values in the imagereconstruction algorithm will automatically compensate for the differentspeeds.

Obviously, the probes that have been described for imaging the heartwould work equally well for imaging abdominal organs and other parts ofthe body such as legs, arms, and neck. In fact, use of receive-onlytransducers in conjunction with a transmit/receive probe would workbetter for abdominal organs because the orientation of the probe set isnot limited by the intercostal spaces formed by the ribs. Whereas thelocations of the acoustic windows to the heart limit the orientation ofthe probe to only a few orientations and it is necessary to rock theprobe to gather three dimensional data, the probes can be used on theabdomen in any orientation presently used. Therefore the probes can beused for real-time 2D scans to duplicate presently accepted proceduresexcept with much higher lateral resolution. In fact, this application ofthe technology may be as important as the application to cardiology(which was our original motivation).

For abdominal scanning it is not necessary to have an elaborate spacingadjustment between the active transmit/receive elements and thereceive-only elements. In fact they could all be mounted together in onerigid probe, either as a linear array or an array with known curvature.Some prior art wide linear arrays exist which insonify tissue by using asmall subset of the total number of elements to transmit and receive abeam perpendicular to the array. Then another partially overlappingsubset of elements is used to transmit and receive another line parallelto the first one, and so on until an entire scan is completed.

However, the same array could be partitioned into an active section plusone or more passive sections where all sections would be used for eachpulse. The active section of elements would be used in transmit as asector scanner sending out beams in a sequence of angular paths. Onreceive, all elements would be treated as independent relativelynondirectional receivers and their outputs would be combined to form ahigh resolution image by the methods taught in this patent.Cross-correlation image matching to account for the variations inultrasound speeds could be done separately for each receive element orfor groups of elements for which the speed corrections would be nearlythe same.

The concept of mounting the active and receive-only elements on a rigidstructure eliminates the necessity for articulating and instrumentingthe spacing between elements thus making practical combined probes to beused for trans-esophageal (TEE), trans-vaginal, and trans-rectalimaging.

A final class of probes would involve putting a receive-only transduceror transducers on the end of a catheter to be inserted in an artery,vein, or urethra while a separated transmit transducer array is appliedto the surface of the skin. The advantage of this approach is that thecatheter could be positioned close to an organ of interest therebyreducing the total transit distance from the transmit transducer to thereceive element and thus higher frequencies could be used for betterresolution. The receive element(s) on the catheter would not have to besteered as it (they) would be relatively omnidirectional.

The Wiener Filter:

The Wiener filter itself is not new, but since it is important for thede-convolution step it will be described briefly here in the context ofthe present invention. The Wiener filter is the mean squared erroroptimal stationary linear filter for images degraded by additive noiseand blurring. Wiener filters are usually applied in the frequencydomain. Given a degraded image I(n.m.), one takes the discrete FourierTransform (DFT) or the Fast Fourier Transform (FFT) to obtain I(u,v).The true image spectrum is estimated by taking the product of I(u,v)with the Wiener filter G(u,v):

Ŝ=G(u,v)I(u,v)

The inverse DFT or FFT is then used to obtain the image estimate s(n,m)from its spectrum. The Wiener filter is defined in terms of thefollowing spectra:

(a) H(u,v)—Fourier transform of the point spread function (psf);

(b) P_(s)(u,v)—Power spectrum of the signal process, obtained by takingthe Fourier transform of the signal autocorrelation;

(c) P_(n)(u,v)—Power spectrum of the noise process, obtained by takingthe Fourier transform of the noise autocorrelation;

The Wiener filter is:

${G( {u,v} )} = \frac{H*( {u,v} ){P_{S}( {u,v} )}}{{{{H( {u,v} )}}^{2}{P_{S}( {u,v} )}} + {P_{n}( {u,v} )}}$

The ratio P_(s)/P_(n) can be interpreted as signal-to-noise ratio. Atfrequencies with high signal to noise ratio, the Wiener filter becomesH⁻¹(u,v), the inverse filter for the psf. At frequencies for which thesignal to noise ratio is low, the Wiener filter tends to 0 and blocksthem out.

P_(s)(u,v)+P_(n)(u,v)=|I(u,v)|². The right hand function is easy tocompute from the Fourier transform of the observed data. P_(n)(u,v) isoften assumed to be constant over (u,v). It is then subtracted from thetotal to yield P_(s)(u,v).

The psf can be measured by observing a wire phantom in a tank using theultrasound instrument. The Fourier transform of the psf can then bestored for later use in the Wiener filter when examining patients.

Because the psf is not constant as a function of range, the Wienerfilter will have to be applied separately for several range zones andthe resulting images will have to be pieced together to form one imagefor display. A useful compromise might be to optimize the Wiener filterjust for the range of the object of interest such as a coronary arteryor valve. It will be necessary to store separate Wiener filters for eachomni-directional probe and for the main probe when it is used as areceive transducer.

An alternative to the Wiener Filter for deconvolution is the least meansquare (LMS) adaptive filter described in U.S. patent application Ser.No. 11/532,013, now U.S. Pat. No. 8,105,239. LMS Filtering is used inthe spatial domain rather than the frequency domain, and can be appliedto the radial scan line data, the lateral data at each depth, or bothtogether.

Image sharpening can be accomplished by the use of unsharp masking.Because aperture blur is much more pronounced in the lateral dimensionperpendicular to the insonifying beam) than in the radial dimension, itis necessary to perform unsharp masking in only one dimension. Whenusing a sector scanner, this masking should be performed before scanconversion. When using a linear phased array, the unsharp masking shouldbe performed on each data set of constant range. Unsharp maskingconsists of intentionally blurring an image, subtracting the result fromthe original image, multiplying the difference by an arbitrary factor,and adding this to the original image. In one dimension this is the sameas blurring a line of data, subtracting it from the original line, andadding a multiple of the difference to the original line.

Multiple Active Transducers—Two Alternative Approaches:

It is possible to use more than one active transducer placed at multipleacoustic windows in order to achieve the same goals of increased lateralresolution and noise suppression. A practical method of providingmultiple omni probes is to use a second phased array head in a secondacoustic window and then treating each element or group of elements ofthe second phased array as a separate omni. With this configuration ofprobes it would be possible to switch the functions of the two probeheads on alternate scans thereby generating images with differentspeckle patterns which can be averaged out.

Multiple phased array heads can also be used together so that both areactive on the same scan. When two (or more) phased array transducers areplaced in the same scan plane, they can be programmed with delays suchthat they act as a single array with a gap in the array of transducerelements. The advantages of having a gap in the array include a)achieving the lateral resolution of a wide aperture without the expenseof filling in acoustic elements through the gap, and b) the gap in theprobe or between probes can be fitted over ribs or the sternum. Thefirst advantage applies equally to applications other than cardiac. Thedisadvantages of multiple active probes is that both the transmit andreceive delays have to be recomputed for each new gap dimension and/orangular orientation of one probe relative to the others.

An active probe with a gap has been demonstrated to produce lateralresolution as good as the probe without the gap. This implies thatlarger gaps will achieve higher resolution since lateral resolution isdetermined primarily by the overall aperture. Referring now to FIG. 10A,there is illustrated the image 800 of an ATS Laboratories Model 539phantom was imaged using an Acuson 128 XP-10 ultrasonic scanner with a4V2c probe. In FIG. 10B, the same probe was used to image the phantomwith its center 64 elements totally obscured by aluminum foil andelectrical tape. As can be seen, the lateral resolution in the image 900is as good as the original although the image quality is degraded byspeckle and other noise.

FIG. 11 shows an image 1000 of the same phantom produced by the sametransducer with the center obscured, but with substantial imageprocessing over multiple scans.

FIGS. 12A-E are various views showing an adjustable, extendable handheld two-aperture probe 1100 adapted for use in cardiology US imaging.This apparatus embodies the inventive concept of separating theinsonifying probe 1110 (a transmit transducer) from the imaging elements1120 (receiver transducer). This comfortable device includes adjustmentmeans 1130, such as a scroll wheel, which selectively drives theelements either closer or further apart along either a medialtelescoping portion 1140 or a medial insertable sleeve, and therebyprovides a range of separation at predetermined distances. The grippingportion 1150 provides easy access to the scroll wheel and places theuser's hand in the functional position to minimize overuse injury. FIG.12A shows the probe in a partially extended configuration. FIG. 12Bshows the probe in a collapsed configuration. FIG. 12C shows the probeextended so as to place the heads at a maximum separation distancepermitted under the probe design. FIG. 12D shows the probe in acollapsed configuration, with adjustment means shown. And FIG. 12E is adetailed perspective view showing surface features at the grippingportion of the probe.

FIG. 13 shows a three-dimensional display 1200 of the anechoic tubes ofthe Model 539 phantom. This 3D display was formed from 13 parallelslices produced with the same transducer with the center obscured. Whenthe total aperture is increased it will be possible to display smalleranechoic tubes such as the coronary arteries. The processing involvedfor this display is a combination of the techniques of the instantpatent and those of U.S. patent application Ser. No. 11/532,013 now U.S.Pat. No. 8,105,239.

Having fully described several embodiments of the present invention,many other equivalents and alternative embodiments will be apparent tothose skilled in the art. These and other equivalents and alternativesare intended to be included within the scope of the present invention.

What is claimed is:
 1. A method of ultrasound imaging, comprising:transmitting a first ultrasound signal from a transmit transducer arrayinto a target object; receiving first echoes of the first ultrasoundsignal with a first transducer that is omnidirectional in an image planeand is located in a first acoustic window; storing the first echoes in acomputer memory; receiving second echoes of the first ultrasound signalwith a second transducer that is omnidirectional in the image plane andis located in a second acoustic window that does not overlap the firstacoustic window; storing the second echoes in the computer memory;retrieving the first echoes from the computer memory and constructing afirst image of the target object from the retrieved first echoes;retrieving the second echoes from the computer memory and constructing asecond image of the target object from the retrieved second echoes; andcombining the first and second images to form a combined image.
 2. Themethod of claim 1, further comprising, prior to combining the first andsecond images, determining a displacement of the second image relativeto the first image.
 3. The method of claim 1, further comprisingapplying a Wiener filter to the first and second images before combiningthe first and second images.
 4. The method of claim 1, furthercomprising applying a Wiener filter to the combined image.
 5. The methodof claim 1, wherein the second receive transducer is at least onetransducer element on the transmit transducer.
 6. The method of claim 1,wherein the second transducer is separate from the transmit transducerand the first transducer.
 7. The method of claim 1, wherein the targetobject is human tissue.
 8. The method of claim 7, wherein the firstacoustic window is selected from the group consisting of a firstparasternal intercostal space, a second parasternal intercostal space, asuprasternal notch, a substernal position, a subcostal window, an apicalview, a first intercostal space adjacent the sternum, and a secondintercostal space adjacent the sternum, and wherein the second acousticwindow is selected from the same group.